In conventional gamma cameras such as Anger cameras known in the art, a single radiation detector having a planar surface is employed for detecting gamma rays for tomographic imaging. A radiopharmaceutical or radioisotope, chosen for its affinity for a particular region or metabolic function of interest, is administered to the patient. The radioisotope emits gamma radiation in all directions from the target location or process. Some of the emitted gamma rays leave the body in the direction of the detectors carrying with them information about their location of origin. The sensitivity of conventional gamma cameras to gamma radiation can be increased by employing a multiplicity of detector heads. When performing Single Photon Emission Computed Tomography (SPECT) using such multiple-detector systems, the detectors are caused to orbit the patient or the object of interest in order to sample from many locations around the object the distribution of radioactivity being emitted from the patient or the object of interest. In conventional SPECT using large detectors, the orbiting of the detector or detectors and the sampling of the gamma rays from multiple directions is necessary in order to provide sufficient information to reconstruct a three-dimensional image of the radiation source by means of computed tomography.
In the prior art, the detectors are typically mounted on a gantry to provide structural support and to orbit the detector around the object of interest. The detector is shielded to prevent stray gamma rays (those not originating from the object of interest) from being detected. Between the radiation detector and the object being imaged is a collimator that is used to restrict the acceptance, or the direction of travel, of incident gamma rays. Typically this collimator is constructed to provide a multiplicity of small holes in a dense, high-atomic-number material such as lead. The gamma rays will pass through the holes if they travel in a direction aligned with the hole but will tend to be absorbed by the collimator material if they travel in a direction not aligned with the holes.
Conventional gamma camera detectors used for medical imaging have a large, flat field of view (typically about 300 in.sup.2) and are very heavy, typically weighing several hundred pounds. These detectors must be made to orbit the patient or object of interest as close as possible (for best image quality) and with a high degree of accuracy and precision. In the current state of the art, large gantries and powerful motors are required to control and accomplish this motion. The safety of the patient in such systems is always a concern. The size and weight of conventional systems often limit or preclude entirely certain dynamic (or "real-time") imaging procedures that require quickly obtaining views of the patient from sufficiently many directions to allow multiple time-sequenced tomographic images to be produced. Due to their weight and the requirements for accurate detector rotation, conventional tomographic gamma camera systems impose elaborate site requirements and operating conditions and are practically incapable of being relocated to another site. Thus, there is a need for tomographic systems that are mobile, allowing the imaging system to come to the patient rather than forcing the patient to be transported to the imaging system. Often, it is unsafe, inconvenient or impossible to move the patient to the conventional gamma camera system. System mobility can improve diagnostic cost efficiencies and provide better health care as a result of timely, on-the-spot prognosis. Moreover, the economics of medical imaging would be improved if the weight of the system could be reduced and the site conditions for the system could be simplified. Thus, there exists a clinical need for tomographic systems that are smaller, that avail themselves to existing clinical applications as well as emerging special-purpose clinical applications, and which impose less stringent site requirements.
Many of the current problems in nuclear medicine imaging are caused by the large flat face of the conventional detector. The algorithms used to determine the location of the gamma interaction (the gamma event) in Anger cameras favor large, flat crystals. These positioning algorithms break down near the edges of the detector producing a significant "dead margin" of several centimeters around the perimeter of the detector. Not only is the dead margin unable to produce usable information, it makes difficult or impossible several types of clinical acquisition protocols, such as lateral breast imaging adjacent to the chest wall and SPECT imaging of the breast.
One of the major problems in nuclear medicine is the poor spatial resolution in the images. This poor spatial resolution is caused mostly by the collimator, and only slightly by the intrinsic resolution in the detector. The collimator restricts the angle of acceptance of incident gamma rays and thus produces a distance-dependent resolution that grows linearly with distance between the source and detector. Roughly speaking, there is loss of 1 mm in resolution for every 1 cm distance between the source and the detector. Simple geometry and the need to clear the patient during an orbital scan prohibit the entire camera face from approaching the patient. Inevitably, in nearly all SPECT applications, part of the camera face is far from the patient and suffers serious loss of resolution.
One clinical application that can benefit from improved spatial resolution is the detection and diagnosis of breast cancer. X-ray Mammography is the standard procedure used in detecting small non-palpable abnormalities in breast tissue. Although modern x-ray mammography has a sensitivity of nearly 90% for the detection of breast cancers, its specificity (the ability to distinguish malignant from benign tissue) is rather low. Of all the breast biopsies performed because of suspicious x-ray mammograms, only about 11-36% are positive for cancer. The path to final diagnosis, which may typically involve percutaneous fine-needle aspiration, stereotactically guided core biopsy, or surgical excision, is often long, expensive, and emotionally and physically traumatic for the patient. Nuclear Mammography is a promising new technique for the detection of breast cancer. Nuclear Medicine studies are a unique and valuable clinical tool. They can distinguish benign from malignant lesions based on cell metabolism, providing a non-invasive cost-effective intermediate option before resorting to biopsy. In recent years, the importance of Nuclear Mammography and the number of clinical studies performed has grown rapidly.
Nuclear Mammography has been shown to be a promising alternative in the process of locating and diagnosing larger breast lesions. However, Nuclear Mammography with conventional gamma cameras suffers from several drawbacks which severely limit its utility in breast imaging applications. Conventional detectors and collimators lack sufficient spatial resolution to detect and image lesions smaller than about 10 mm. The insensitive "dead margin" near the edge of the detector keeps the useful field of view several centimeters away from the chest wall and limits the amount of the breast that can be imaged laterally. The large size and weight of the detector often limit the clinical applications to lateral planar imaging and preclude medial breast imaging and conventional orbital SPECT of the breast. Tomography would be preferred over planar imaging because SPECT generally provides better lesion contrast.
A need exists for a gamma camera able to perform tomographic imaging without using large detectors or collimators or the orbiting motion of such detectors or collimators. A non-orbiting tomography system composed of smaller detector modules would enable improvements spatial resolution and system sensitivity which would result in improved image quality and diagnostic accuracy. It would also broaden clinical utility, improve patient access to systems, allow better dynamic studies and reduce site requirements.
Several devices or inventions have been made which try to address some of these issues. Wong and Hicks (U.S. Pat. No. 5,451,789), for example, disclose an invention in which uncollimated, stationary detectors perform breast tomography by means of positron emission (dual-photon) tomography (PET). In contrast, the invention disclosed here combines collimation, localized (non-orbiting) detector or module motions, and various module/acquisition geometries to perform tomography of the breast and other objects of interest by means of single-photon emission tomography. Similarly, the system described by Genna et al (U.S. Pat. No. 4,584,478) is based on a collimator which must orbit the object of interest in order to perform tomography. The system described by Ashburn (U.S. Pat. No. 5,742,060), while mobile, is not capable of tomography. The Headtome II device, (as described for instance in "A Hybrid emission CT-Headtome II", by Y. Hirose et al, IEEE Transactions on Nuclear Science, Vol. NS-29, No. 1, February, 1982, pp. 520-523), employs orbital and "wobbling" motions of its detector and collimator. However, the wobbling motions are not sufficient for tomography and must be augmented by orbiting detector or collimator motions. An important component of the invention disclosed here is the non-orbital motion of the detectors or collimators which are sufficient for tomography.